The present invention generally relates to ultrasonic probes and more specifically to ultrasonic probes for acoustic imaging.
Ultrasonic probes provide a convenient and accurate way of gathering information about various structures of interest within a body being analyzed. In general, the various structures of interest have acoustic impedances that are different than an acoustic impedance of a medium of the body surrounding the structures. In operation, such ultrasonic probes generate a beam of broadband acoustic waves that is then coupled from the probe, though a lens, and into the medium of the body so that the acoustic beam is focussed by the lens and transmitted into the body. As the focussed acoustic beam propagates through the body, part of the signal is reflected by the various structures within the body and then received by the ultrasonic probe. By analyzing a relative temporal delay and intensity of the reflected acoustic waves received by the probe, a spaced relation of the various structures within the body and qualities related to the acoustic impedance of the structures can be extrapolated from the reflected signal.
For example, medical ultrasonic probes provide a convenient and accurate way for a physician to collect imaging data of various anatomical parts, such as heart tissue or fetal tissue structures within a body of a patient. In general, the heart or fetal tissues of interest have acoustic impedances that are different than an acoustic impedance of a fluid medium of the body surrounding the tissue structures. In operation, such a medical probe generates a beam of broadband acoustic waves that is then acoustically coupled from a front portion of the probe, through an acoustic lens, and into the medium of the patient's body, so that the beam is focussed and transmitted into the patient's body. Typically, this acoustic coupling is achieved by pressing the front portion of the probe having the lens mounted thereon into contact with a surface of the abdomen of the patient. Alternatively, more invasive means are used, such as inserting the front portion of the probe into the body through a catheter.
As the acoustic signal propagates through the patient's body, part of the acoustic beam is weakly reflected by the various tissue structures within the body and received by the front portion of the ultrasonic medical probe. By analyzing a relative temporal delay and intensity of the weakly reflected waves, an imaging system extrapolates an image from the weakly reflected waves. The extrapolated image illustrates spaced relation of the various tissue structures within the patient's body and qualities related to the acoustic impedance of the tissue structures. The physician views the extrapolated image on a display device coupled to the imaging system.
Because the acoustic beam produced by these ultrasonic probes is only weakly reflected by the tissue structures of Interest, it is important to try to concentrate the acoustic beam by efficiently focussing the acoustic beam. Such efficient focussing would insure that strength of the acoustic beam generated by the probe is enhanced as the signal is transmitted from the front portion of the probe, through the lens, and into the medium of the body. Additionally, such efficient acoustic focussing would insure that the weakly reflected acoustic waves are concentrated as they pass though the lens to be received by the front portion of the probe. Focussing is also desired to provide improved imaging resolution of structures within the body under examination.
Furthermore, since the acoustic waves are only weakly reflected by the tissue structures of interest, it is important to reduce any extraneous acoustic signals transmitted or received by the probe through the acoustic lens. In general, any physically realizable acoustic radiator has some finite aperture. As representatively illustrated in FIG. 1, diffraction of the acoustic waves 101 through the finite aperture, E, results in a desired main lobe 105 and undesired side lobes 107 arranged in a familiar intensity pattern corresponding to a function (sin x)/x. For example, if the acoustic beam generated by the probe is diffracted through the finite aperture of the acoustic lens, then a desired acoustic signal is transmitted into the patient along a main transmission lobe of the beam, and a first extraneous acoustic signal is transmitted into the patient along side transmission lobes of the beam. Similarly, because of the finite aperture of the acoustic lens, the probe receives another extraneous acoustic signal along side reception lobes in addition to reflected acoustic waves along the main reception lobe. Such extraneous acoustic signals can distort the extrapolated image viewed by the physician unless corrective measures are undertaken.
As previously known an ultrasonic probe comprises a layer of a dissimilar acoustic material adhesively bonded to a rear portion of a piezoelectric vibrator body. A thin layer of a cement adhesive is applied to bond each layer, thereby creating undesirable adhesive bond lines between the layers of dissimilar material and the piezoelectric body. The layer of material is in turn coupled to the acoustically damping support body. For example, FIG. 2 illustrates an ultrasonic transducer 200 comprising a piezoelectric vibrator body 204 of a piezoceramic, such as lead zirconate titanate having the acoustic impedance of 33*10.sup.6 kg/m.sup.2 s, a layer of dissimilar acoustic material such as silicon 206 having an acoustic impedance of 19.5*10.sup.6 kg/m.sup.2 s, a support body 208 of epoxy resin having an acoustic impedance of 3*10.sup.6 kilograms/meter.sup.2 second, kg/m.sup.2 s. The vibrator body 104 shown in FIG. 1 has a resonant frequency of 20 megahertz, MHz, and the silicon layer has a thickness that is a quarter wave length of the resonant frequency of the vibrator body. Electrodes 210 are electrically coupled to the vibrator body 204 for electrically sensing acoustic signals received by the transducer.
The piezoelectric vibrator body 204 shown in FIG. 2 is connected on one side to the silicon layer by means of an adhesive layer 212. The thickness of the adhesive layer is typically 2 microns. A silicon layer adhesively bonded to a piezoelectric body is also discussed in U.S. Pat. No. 4,672,591 entitled "Ultrasonic Transducer" and issued to Briesmesser et al. Because this patent provides helpful background information concerning dissimilar acoustic materials bonded to piezoelectric bodies, it is incorporated herein by reference.
Though the dissimilar acoustic matching materials employed in previously known schemes provides some advantages, the adhesive bonding of these layers creates numerous other problems. Bonding process steps needed to implement such schemes create manufacturing difficulties. For example, during manufacturing it is difficult to Insure that no voids or air pockets are introduced to the adhesive to Impair operation of the probe. Furthermore, reliability of this previously known transducers is adversely effected by differing thermal expansion coefficients of the layers of dissimilar materials and the piezoelectric ceramic bodies. Over time, for example over 5 years of use, some of the adhesive bonds may lose integrity, resulting in transducer elements that do not have efficient acoustic coupling to the damping support body. Additionally, operational performance is limited at higher acoustic signal frequencies, such as frequencies above 20 megahertz, by the bond lines between the piezoelectric body and the dissimilar materials.
One measure of such operational performance limitations is protracted ring down time in impulse response of the ultrasonic transducer of FIG. 2. Such impulse response can be simulated using a digital computer and the KLM model as discussed in "Acoustic Waves" by G. S. Kino on pages 41-45, which is incorporated herein by reference. FIG. 3 is a diagram of the simulated impulse response of the ultrasonic transducer of FIG. 3 having the resonant frequency of 20 Megahertz, radiating into water, and constructed in accordance with the principles taught by Briesmesser et al. In accordance with the impulse response diagram shown in FIG. 3, simulation predicts a -6 decibel ,db, ring down time of 0.221 microseconds, usec, a -20 db ring down time of 0.589 usec, and a -40 db ring down time of 1.013 usec.
Another previously known ultrasonic probe includes high-polymer piezoelectric elements. Each of the high-polymer piezoelectric elements comprises a composite block of piezoelectric and polymer materials. Such composites are discussed in U.S. Pat. No. 5,142,187 entitled "Piezoelectric Composite Transducer For Use in Ultrasonic Probe" and issued to Saito et al. Because this patent provides helpful background information concerning piezoelectric composites, it is incorporated herein by reference.
While composite materials provide other advantages, there are difficulties in electrically sensing reflected acoustic waves received by such composites. A dielectric constant of each high polymer element is relatively small. For example, for a composite that is 50% polymer and 50% piezoelectric ceramic, the dielectric constant measurable between electrodes of the high polymer element is approximately half of that which is inherent to the piezoelectric ceramic. Accordingly, the dielectric constant measurable between the electrodes of the high polymer element is only approximately 1700. A much higher dielectric constant is desirable so that a higher capacitive charging is sensed by the electrodes in response to the reflected acoustic waves. The higher dielectric constant would also provide an improved electrical impedance match between the probe and components of the imaging system electrically coupled to the probe.
What is needed is a reliable ultrasonic probe that provides enhanced operational performance, efficient electrical coupling to imaging system components, focussing of the main lobe of the acoustic beam, and reduced side lobes.